This invention relates to a charge detector for reading charge produced by an active pixel in the detector.
A known diagnostic technique used in tomography for locating tumors involves injecting into a patient""s bloodstream a radioactive isotope which targets the tumor, so that the location of the tumor can be derived by detecting the location of the radioactive isotope. Typically, the radioactive isotope emits xcex3-rays which are dispersed from the tumor site. In order to achieve the desired detection so as to determine the precise location of the tumor, it is necessary to image the patient""s body in such a manner as to detect only those xcex3-rays which are emitted normally from the body and to ignore those xcex3-rays which are dispersed in other directions.
U.S. Pat. No. 5,656,818 (Nygard) assigned to the one of present applicants discloses such a radiation imaging system that includes a detector unit and a receiver unit. The detector unit includes a two-dimensional sensor, first and second amplifier channels, first and second multiplicity generators and first and second address generators. The two-dimensional sensor includes first and second sensing segments that sense radiation in a first direction and a second direction, respectively. Each first amplifier channel generates an output signal based on a detection output from a corresponding one of the first sensing segments. The first multiplicity generator generates a first multiplicity signal representing a number of the first amplifier channels generating output signals. The first address generator generates a first analog address of a first amplifier channel associated with a received output signal. The second amplifier channels, second multiplicity generator and second address generator operate in like manner with respect to the second sensing segments. The receiver unit includes converters for converting the first and second analog addresses into first and second digital addresses. The receiver unit also includes a tester for testing whether the first and second digital addresses represent a valid position address in the first and second direction based on the first and second multiplicity signals.
Different types of computer tomography are known in which such a radiation imaging system may be embodied. In Single Photon Emission Computed Tomography (SPECT) more than one detector is rotated around the subject. During the rotation of the detector, the counting of the gamma rays is repeated. Then, the radioisotope""s distribution (tomographic image) is reconstructed based on the obtained count values of the xcex3-rays.
In contrast to SPECT where a radioisotope in the body emits xcex3-rays produced by a single photon, in Positron Emission Tomography (PET) a patient is administered a radioisotope that emits positrons (i.e. positively charged electrons). When the positrons meet electrons within the body, the positrons and electrons mutually annihilate and produce two xcex3-rays that propagate away from each other at an angle of 180xc2x0 and are detected by respective detector segments in the PET scanner. The scanner""s readout electronics record the detected xcex3-rays and map an image of the area where the radioisotope is located. Here also two simultaneous detections are indicative of a positron emission from the tumor site.
Thus, in PET two simultaneous xcex3-rays must be detected on opposite sides of the patient""s body. The positrons are extremely short lived and simultaneity implies that the two xcex3-rays can both be detected within a short time window, which is 10 ns, for example. The PET scanner surrounds the patient like a CT scanner and typically comprises in the order of 200,000 pixels on two detector segments. Thus, it is necessary to detect two excited pixels within a short time difference and then to read out the energy of these two pixels. Only if the energy of each active pixel equals about 511 keV, (i.e. the energy of the incident xcex3-radiation) are the two photons the result of positron-electron annihilation and thus indicative of the tumor""s location. Also in a PET scanner Compton scattering can occur within the detector, whereby a photon is only partially absorbed by the pixel and partially scattered to another pixel or occasionally to even more than a single pixel. In this case the sum of the energies of simultaneous active pixels will be equal to about 511 keV.
Another apparatus used for nuclear imaging is the Compton Camera. In the Compton Camera in order to determine the location of the xcex3-ray based on the detection of the single photon thus emitted, it is necessary that Compton scattering occur, so that another photon will be emitted substantially simultaneously, thereby allowing the angle of the incident xcex3-ray to be calculated.
Here, too, it is necessary to establish simultaneity of a xcex3-ray striking multiple pixels, although Compton Camera and PET must detect simultaneous xcex3-rays and so the time difference for establishing simultaneity of two or more pixels becoming active is more critical than SPECT. Known architectures for PET scanners use sequential readout whereby each pixel is read sequentially in order to determine whether or not it is active.
In EP 893 705 published on Jan. 27, 1999 entitled xe2x80x9cMulti-Channel Reading Circuit for Particle Detectorxe2x80x9d and assigned to one of the present applicants, there is described a method for reading an array of pixels in a 2-dimensional image sensor so as to reduce the time taken to detect a single xe2x80x9cactivexe2x80x9d pixel. Specifically, it is proposed in EP 893 705 to group pixels into predefined groups or segments (typically associated with a single ASIC) and to read out the pixels of only those groups of pixels that are found to contain an active pixel. Such an approach is based on the principle that determination of an active group may be made quickly providing that, at this initial stage, it is not required to determine which specific pixel within the active group is itself active. Having thus established which groups of pixels are active, each pixel in the active group is then read sequentially.
Whilst this is suitable for SPECT, it is not suitable for PET principally due to the very high rate at which pixels are activated in a PET scanner owing to its absence of a collimator. As a result, there are too many active pixels to make it practical to provide a tradeoff between providing a large number of groups or segments each containing a sufficiently small number of pixels that sequential readout in each group can be achieved in reasonable time. Specifically, whilst it is theoretically possible to reduce the number of pixels in each group so that pixels in each group that become active simultaneously within the 10 ns time window can be determined, the number of A/D converters that would then be required to effect digital processing of the active pixels would, in practice, be prohibitive.
The reading circuit described in EP 893 705 uses two detectors so as to detect Compton scattering, requiring that data be read out in the first detector from a large number of addressable pixels along respective channels in order to detect which pixel is xe2x80x9cactivexe2x80x9d. This is done by first integrating the charge associated with each pixel using an integrator in the form of an operational transconductance amplifier having a feedback capacitor. The integrated charge pulse is then amplified and shaped and the resulting analog signal is sampled and held, allowing its magnitude to be measured. In order to measure the peak magnitude of the shaped signal, the shaped signal must be very accurately sampled at the peak value. This requires an accurate determination of the peak time that occurs a fixed time difference tp after the emission of charge by the excited pixel. The fixed time difference tp is a function of the RC time constant of the shaper circuit and is therefore known.
Thus, in order to know when to sample the integrated charge signal, the time of occurrence to of each charge emission must itself be accurately determined. This having been done, all that is then necessary is to sample the held integrated charge sample at a time tp. A reading system for reading out the charge signals must therefore generate an accurate trigger simultaneous with the occurrence of each charge emission.
Obviously, if during every scan of the composite image sensor, each pixel is read sequentially only one at a time, then the current scan in each segment can be terminated when an xe2x80x9cactivexe2x80x9d pixel is detected assuming that only pixel in each segment can be active. However, it is impractical to read each pixel in such a manner because of the time overhead involved in addressing each pixel separately and downloading the pixel data along a dedicated channel for further processing.
The need to avoid sequential readout is addressed by U.S. Pat. No. 5,847,396 (Lingren et al.) assigned to Digirad Corporation, which discloses a high-energy photon imaging system comprising an imaging head that includes a detector having a plurality of detection modules. Each detection module comprises a plurality of detection elements fixed to a circuit carrier. The detection elements produce electrical pulses having amplitudes indicative of the magnitude of radiation absorbed by the detection elements. The circuit carrier includes channels for conditioning and processing the signals generated by corresponding detection elements. Each conditioning and processing channel stores the amplitudes of the detection element electrical pulses exceeding a predetermined threshold. The detection modules employ a fall-through circuit, which avoids the need for sequential readout and automatically finds only those detection elements whose stored pulse amplitude exceeds the threshold. The fall-through circuit searches for the next detection element and associated channel having a valid event, meaning that the detection element exhibits a pulse magnitude that exceeds a certain threshold.
Lingren et al. is directed to a xcex3-ray camera and imaging system where both planar and SPECT images may be obtained. It is not directed to a PET system. The distinction is significant. SPECT systems, by definition, produce only a single emission from each xcex3-ray. Tomography is used to determine from multiple emissions, each deriving from a different xcex3-ray striking the detector, the source of emission of the xcex3-rays, i.e. the location of the tumor. Since each xcex3-ray derives from only a single emission, all such emissions are valid, providing only that the corresponding pulse magnitude consequent to the emission exceeds the threshold. This is determined using a threshold comparator coupled to each detection element, such that the output of each threshold comparator establishes whether or not the event is xe2x80x9cvalidxe2x80x9d. It is important to note that the actual energy level of the detection element is unimportant: all that is relevant is that its energy level exceeds the threshold. Thus, once an event is established as being valid, the fall-through circuit merely ripples through the valid events quickly in order to determine which detection elements are xe2x80x9cactivexe2x80x9d. Specifically, once a detection element is established by the threshold comparator as being valid, no further information regarding that detection element is required in order to effect the subsequent stage of computer tomography.
All this is very far from the case in PET systems where the mere establishment that a pixel is active does not mean that the event giving rise to that pixel becoming active is xe2x80x9cvalidxe2x80x9d. In PET systems the time of emission is of vital importance since only those emissions which occur simultaneously are of interest and any pixels that become active even within so short a differential time such as 10 ns for example must be disregarded as PET-type emissions. Furthermore, each xcex3-ray emission deriving from mutual positron-electron annihilation has an energy of 511 KeV and thus the peak energy of each emission is best measured and used to establish validity of the emission as a PET event. Such establishment is further complicated by the fact even a xcex3-ray emission deriving from mutual positron-electron annihilation may give rise to Compton scattering on the detector, in which case the energy of an active pixel may be less than 511 KeV whilst still constituting a valid event since it can be mated with one or more counterpart emissions whose cumulative energy is equal to 511 KeV. Usually the active pixel having lowest energy is assumed to be the primary pixel first hit by the photon and the other pixel or pixels are ignored. For the sake of completeness, it has to be noted that Compton scattering at the detector can also occur between different detector segments of a multi-segment detector, albeit with much reduced probability. It is possible also in this case to sum the energies of simultaneous active pixels between detector segments so as to establish whether active pixels in adjacent detector segments result from Compton scattering at the detector. In fact, it is not essential to consider such emissions since tomography can still be performed using only that pixel information relating to direct PET-type emissions. However, this results in useful information being ignored and this, in turn, means that a greater number of actual direct PET-type emissions must be collected in order to be able to perform tomography accurately. The fact that more direct PET-type emissions must be collected means that the time taken to collect the requisite data is increased and this, of course, means that the patient is exposed to more irradiation, which is undesirable.
It is thus clear that the system described by Lingren et al. does not address the very specific requirements associated with PET systems, where validity of an active pixel cannot be established only on the basis of the energy level exceeding a threshold and simultaneity and actual energy level of the pixel must be established very quickly in order to establish validity.
Simultaneity of xcex3-ray stimulated emissions in two opposite detector segments of a PET camera is established by correlating events in the two detector segments in order to establish that they derive from positron-electron annihilation giving rise to two xcex3-rays. As noted above, this is done by establishing that the two events are substantially simultaneous. Moreover, accurate simultaneity of the two events can be determined accurately only if the xcex3-ray emission is measured fast. Use of a single filter having a slow time constant for shaping the data signal resulting from the xcex3-ray emission, whilst commonly used, detracts from the accuracy with which the peak time can be measured and this, in turn, reduces the accuracy with which simultaneity of corresponding events in two detectors can be established.
EP 893 705 proposes the use of a fast shaper for determining the initiation of a xcex3-ray emission so as to establish when a pixel becomes active, whereupon a slow shaper determines the peak time. As noted above, this is used within the readout circuit of a Compton camera. However, accurate determination of the initiation of a xcex3-ray emission by the fast shaper serves to allow accurate sampling of the energy profile filtered by the slow shaper and sampled and held for subsequent measurement. This is done by reading the energy profile a known time delay after establishing initiation of the xcex3-ray emission by the fast shaper.
It is specifically to be noted that EP 893 705 does not rely on the accurate measurement of time of emission in order to establish simultaneity of emissions. Rather this is determined using logic gates connected to different detector segments so that segments that contain active pixels simultaneously are identified. Associated with each detector are groups of pixels, each group containing a plurality of pixels. A logic level associated with each group of pixels indicates which group contains an active pixel. Thus, simultaneity of two events, deriving from Compton scattering of a single photon, results in two detectors being active and of two groups, one in each active detector, being active: all this being determined using logic gates in real time. Time stamps are associated with each pixel but it is specifically noted that these are not used to establish simultaneity of a charge emission from the first and second detectors. Rather, as is also explained, the time stamp identifies the time of emission of the active pixel for the sole purpose of measuring the peak value of the energy profile filtered by the slow shaper. This is done after establishing simultaneity of active pixels in the manner explained above.
In PET, owing to the fact that the double photon emission derives from mutual annihilation of a single positron with a single electron, only two simultaneous photon emissions are possible. In practice, such occurrence can only be determined retroactively by detecting multiple photon emissions, establishing simultaneity and accepting simultaneous emissions as due to PET only if the number of such emissions is less than a predetermined threshold, typically two. This implies that the actual time of emission must be measured and recorded for subsequent analysis and, of course, that the time must be measured extremely accurately and quickly in order to allow for subsequent accurate matching of simultaneous emissions.
It has further to be borne in mind that use of PET does not prevent the occurrence of other physical phenomena, which have no relevance to PET and must not be allowed to corrupt the simultaneity measurements. Thus, simultaneous emissions do not necessarily indicate positron-electron annihilation since they could be due to Compton scattering within the body of the patient, for example, giving rise to two simultaneous from the body. As against this, however, Compton scattering can occur at the detector. In this case, the multiple photon emissions actually derive from only a single photon emitted from the tumor site, giving rise to a secondary effect when the photon strikes the detector, since not all of its energy is given up. As noted above, Compton scattering results in two (or more) active pixels having a combined energy of 511 KeV allowing it to be determined whether multiple simultaneous events are due to Compton scattering at the detector. In this case they cannot be ignored and usually the active pixel having lower energy is assumed to be the primary pixel first hit by the photon and the other pixel or pixels are ignored. For the sake of completeness, it has to be noted that Compton scattering at the detector can also occur between different detector segments, albeit with much reduced probability. It is possible also in this case to sum the energies of simultaneous active pixels between detector segments so as to establish whether active pixels in adjacent detector segments result from Compton scattering at the detector.
In summary, Lingren et al. disclose the use of fall-through in a SPECT system but is not concerned with the time of a photon""s emission. EP 893 705 discloses the use of a fast shaper for accurately determining an initiation of a photon emission but uses this only to know when to read the energy level as filtered by a slow shaper and held by a sample and hold circuit. Thus, neither of these references addresses the very specific and stringent requirements inherent in PET systems where validity of two simultaneous events amongst a vast number of pixels must be done fast and requires both accurate time stamping of the events and subsequent measurement of the pixel energy.
It is therefore an object of the invention to provide a reading circuit for reading one or more xe2x80x9cactivexe2x80x9d pixels in a 2-dimensional image sensor having a plurality of pixels.
According to a broad aspect of the invention there is provided a method for detecting an active pixel in a sensor segment having a plurality of addressable pixels, the method comprising the steps of:
(a) for each trigger event that causes a pixel to go active, setting a corresponding resettable latch from an initial unset logic state to a set logic state and establishing a time window during which another pixel which becomes active in said sensor is assumed to originate from said trigger event,
(b) for all pixels that become active within said time window:
i) sampling and holding a corresponding analog value of the pixel in response to a logic signal indicative of another pixel currently being active,
ii) establishing an incident time of the pixel going active,
iii) sparsely reading each of the latches in said segment so as successively to identify latches which are in said set logic state each corresponding to an active pixel having a known address in said segment,
iv) for each active pixel associating the known address thereof with the incident time of the pixel going active and the respective sampled and held analog value,
v) repeating steps iii) and iv) in respect of successive active pixels, and
vi) resetting said latches.
According to another aspect of the invention, there is provided a simultaneity detector for detecting simultaneously active pixels in a sensor having at least two addressable segments each containing a plurality of addressable pixels, the simultaneity detector comprising:
a respective sample and hold unit coupled to each pixel in each of the segments, each for sampling and holding a corresponding analog value associated with the pixel,
a respective resettable latch coupled to each of the sample and hold units for changing from an initial unset logic state to a set logic state when the corresponding pixel goes active thereby establishing a time window within which any other pixel that subsequently goes active is considered to be a simultaneously active pixel,
a logic circuit coupled to all segments in the detector for detecting two simultaneous segments in each of which there exists at least one active pixel,
a lookup table in each segment having a plurality of addressable locations each corresponding to a respective pixel in the respective segment and storing the known address of the respective pixel,
an analog multiplexer in each segment having a plurality of addressable channels each coupled to a respective sample and hold unit in said segment for carrying the corresponding sampled and held value of the respective pixel,
a sparse readout circuit in each segment coupled to each of the latches in the respective segment for sparsely reading each of the latches therein so as successively to identify latches which are in said set logic state each corresponding to an active pixel having a known address in the respective segment, said sparse readout circuit being responsive to each of the set latches in turn for pointing successively to an addressable location in said lookup table so as to read the known address of the corresponding active pixel, and for addressing a corresponding channel of the analog multiplexer so as to derive the respective analog value of the corresponding active pixel, and
a reset unit coupled to each of the latches in the respective segment for resetting said latches at a termination of said time window or within said time window if no simultaneous segments are detected.
In accordance with another aspect of the invention, there is provided a method for reading an analog data signal emitted by an active pixel in a sensor having a plurality of addressable pixels, the method comprising the steps of:
(a) converting the analog data signal associated with the active pixel to a digital signal having a first logic state and converting corresponding analog data signals associated with inactive pixels in the sensor to corresponding digital signals each having a second opposite logic state,
(b) using said digital signals to identify the active pixel, and
(c) reading a magnitude of the analog data signal in respect of said active pixel.
According to a preferred embodiment, each of the latches comprises a Flip Flop that is set to HIGH when the corresponding pixel is xe2x80x9cactivexe2x80x9d and otherwise remains LOW. There is thus provided a bank of Flip Flops mapping the pixels in the sensor. However, whilst each of the active pixels in the pixel array has associated therewith an analog energy value, each of the corresponding Flip Flops has associated therewith a digital value HIGH (logic xe2x80x9c1xe2x80x9d) or LOW (logic xe2x80x9c0xe2x80x9d) which is used to address a lookup table in which there is stored an address corresponding to each Flip Flop and thus to each pixel in the sensor. The location of a HIGH Flip Flop in the bank of Flips Flops thus maps the location of the corresponding active pixel in the sensor and can be fed to an N-to-1 analog multiplexer having N inputs corresponding to N pixels and a single output channel for selecting the pixel corresponding to the active input. Thus, in the event of simultaneous emissions from more than one pixel, more than one Flip Flop will be HIGH although the energy values associated with the active pixels is still not known. The energy values of active pixels are determined using a fast shaper to capture the time of the peak and a slow shaper to determine the peak accurately, as described in EP 893 705. However, in order to read out the energy values, it is first necessary to determine which pixels are active and, of course, to do this quickly. In the invention, this is accomplished by using a sparse readout circuit to ripple through the Flip Flops and determine one by one which Flips Flops are HIGH. This can be done extremely quickly and as each HIGH Flip Flop is thus detected, corresponding to an active pixel, the address of the corresponding pixel is read from the lookup table and the corresponding energy value is accessed by the analog multiplexer. Such an approach obviates the need to read the pixels in the non-active groups, thereby saving considerable reading time.
Preferably, the analog value of the pixel is a first rising current pulse derived from an emission of electric charge consequent to the pixel being struck by a xcex3-ray. The current pulse is integrated by a preamplifier so as to produce an analog voltage step having a sharp change in level upon emission of the charge signal. The voltage step constitutes an initiation signal indicative of the time of emission to and whose magnitude is proportional to the accumulated charge produced by the current pulse and which is collected by a feedback capacitor in the preamplifier. The reading circuit further includes at least one shaper in respect of each pixel in the active group which is responsive to the voltage step for amplifying and shaping the integrated charge in order to generate a slowly rising analog voltage signal having a high signal to noise ratio. An important feature of such an embodiment resides in the precision with which the shaped analog voltage signal is samples at its peak. Specifically, the reading circuit includes in respect of each pixel:
a fast shaper having a fast time constant coupled to the respective pixel for producing a timing pulse by fast shaping an output from said pixel, the timing pulse having a known relationship to an incident time of the respective pixel going active,
a slow shaper having a slow time constant coupled to the respective pixel for simultaneously shaping the charge signal associated with the respective pixel so as to generate a slow response curve having high signal to noise ratio,
a threshold discriminator for determining a time delay xcex94t for the fast response curve to exceed said predetermined threshold, and
sampling circuit for sampling the slow response curve at a further time interval tpxe2x88x92xcex94t where tp is the time at which the slow response curve reaches its peak value so as to sample the slow response curve substantially at its peak value.